Positron Emission Tomography (PET) has become a main stream medical imaging technique for applications in oncology, cardiology, neurology, and other medical disciplines. In a PET imaging device, pairs of gamma photons resulting from a positron-electron annihilations within a patient are detected in detectors surrounding the patient. The photons in a given pair travel in opposite directions defining a line of response (LOR), and two detectors on opposite sides of the patient generate signals that represent the respective locations of two points along the LOR. The signal information from the detectors is used to reconstruct a location of the annihilation event.
PET has experienced dramatic growth since 1998 when health insurance providers permitted the reimbursement for using PET imaging. PET has also become a widely-used and critical tool in the areas of biomedical research, drug discovery and development, and so called pre-clinical studies based on molecular imaging. However, PET has faced a fundamental technical challenge with regard to image resolution. The main technical problem is that current PET detectors can only detect two-dimensional signals (an X-Y location), and the Z location coordinate is an approximation based on the location of the detector scintillation crystal. This situation results in poor spatial resolution associated with parallax error due to the crystal penetration by energetic gamma photons. It is a recognized problem that spatial resolution suffers near the radial edges of the field-of-view (FOV), resulting in a lack of spatial resolution uniformity over the entire FOV. PET performance can be significantly improved if PET detectors can measure the Z axis location or so-called “depth-of-interaction” (DOI) of a gamma photon interaction within a detector crystal.
As a result, there have been many suggestions to develop a three-dimensional PET detector over the past decades. The most promising solution so far is to use two photon sensors, such as two avalanche photo diodes (APDs), coupled to measure the interactions from two opposite ends of a detector crystal instead of a single photon sensor from one end according to current two-dimensional detector design.
Although this new design has made significant progress, its widespread adoption may take a few more years due mainly to cost issues associated with photon sensors. But, all major vendors and leading research institutes (including Lawrence Berkeley National Laboratory) are pushing this new technology. As photon sensor technology matures and new photon sensor fabrication processes are developed, improved photon sensors are becoming available at a lower cost, and the commercial application of this three-dimensional design can be expected in the next generation of PET scanners.
With this new PET detector design, there is one major technical challenge: it needs an efficient and accurate calibration method to determine the response function of the three-dimensional detector. Such an efficient and accurate calibration method is crucial to ensure proper functioning of the detector. Since a ratio of the signals from the two photon detectors is related to the DOI along the crystal Z axis, the DOI can be calculated based on this signal ratio if a predetermined function between the signal ratio and DOI is known. This DOI function needs to be experimentally determined through calibration. Currently, this DOI function is determined by using a point radiation source P in the middle of the primary detector crystal C (with its DOI function to be determined) and a second small size detector F, as shown in FIG. 1. A coincidence between the two detectors C and F is required to take the data so that the DOI position can be allocated. Signals are acquired from the two photon detectors APD1 and APD2, and a corresponding signal ratio is calculated for the associated DOI position. This procedure is repeated for a plurality of DOI positions along the depth (Z) axis of detector crystal C to obtain the functional relationship (DOI function) between DOI position and signal ratio. In order to measure the entire DOI function, point source P and second detector F must be moved along the depth axis of primary detector C to take multiple measurements at different DOI positions. There are several drawbacks with this method, including that it requires complicated experimental setup with an extra detector and coincidence process to allocate the DOI position; the measurement has to be taken at different DOI positions along the crystal, which requires demanding detector/source alignment and movement and lengthy time to complete; and it is potentially prone to significant measurement error due to the setup itself and size limitations of the second detector F. Moreover, the current method is simply not practical to measure the DOI functions for multiple crystals inside an array of crystals.
At present, there is no effective calibration method for such three-dimensional PET detectors. The current method described in the preceding paragraph requires a movable radiation source to be accurately aligned with the detector, which is a complicated setting, and cannot be applied to a practical detector since the inner crystals in an array cannot be calibrated. The calibration process is very long and tedious, and is not useful for practical purposes. In summary, the current method of determining a DOI function requires complicated settings for laboratory bench-top testing, and still cannot be applied to any practical detectors, or is too coarse or unstable to be used for accurate calibration required by medical industry standards.